Method and apparatus for determining respiratory system resistance during assisted ventilation

ABSTRACT

Method and apparatus are described for determining respiratory system resistance (R) in a patient receiving gas from a ventilator. A negative pulse in the pressure and/or flow output of the ventilator during selected inflation cycles is generated and Paw, {dot over (V)} and V are measured at a point (T o ) near the beginning of the pulse, at a point (T 1 ) near the trough of the negative pulse and at a point (T −1 ) preceding T o . The value of R is calculated from the difference between Paw, {dot over (V)} and V at T o  and at T 1  and where the change in patient generated pressure (Pmus) in the interval T o −T 1  is estimated by extrapolation from the different between Paw, {dot over (V)} and V and T o  and at T −1 , in accordance with Equation 8.

REFERENCE TO RELATION APPLICATIONS

This application is a U.S. National Phase filing under 35 USC 371 ofPCT/CA01/00578 filed Apr. 25, 2001 claiming priority under 35 USC 119(e)from U.S. Provisional Patent Application No. 60/199,824 filed Apr. 26,2000.

FIELD OF INVENTION

This invention relates to mechanical ventilation, and in particular, toassisted ventilation and the determination of respiratory systemresistance.

BACKGROUND TO THE INVENTION

There are currently no reliable, clinically available, non-invasivemeans to estimate respiratory resistance (R) during inspiration inmechanically ventilated patients who have spontaneous respiratoryefforts. Calculation of resistance requires knowledge of the forceapplied to the respiratory system which, in such patients, includes acomponent related to pressure generated by respiratory muscles (Pmus).This component continuously changes during the inflation phase andcannot be estimated without prior knowledge of respiratory mechanics.Furthermore, to isolate the component of total applied pressure that isdissipated against resistance (P_(res)), it is necessary to subtract thepressure used against the elastic recoil of the respiratory system. Thisrequires knowledge of passive respiratory elastance (E) which is alsodifficult to determine in the presence of unquantifiable Pmus. Atpresent, therefore, R can be reliably estimated only by use ofesophageal catheters, which add another invasive intervention to alreadymuch instrumented patients, or by elimination of respiratory musclepressure output with paralysis, or hyperventilation (controlledmechanical ventilation, CMV). The latter entails additional personneltime and does not lend itself to frequent determination of R. To theextent that R is a highly dynamic property that may change frequently,due to secretions or changes in bronchomotor tone, availability ofcontinuous estimates of R may be helpful in the clinical management ofsuch patients. Thus, changes in R can be rapidly identified and dealtwith. Furthermore, this information makes it possible to adjust thelevel of assist according to the prevailing R values, a feature that isof particular utility in pressure assisted modalities of ventilatorysupport (Pressure Support Ventilation, Proportional Assist Ventilation).

In U.S. Pat. No. 5,884,622 (Younes), assigned to the assignee hereof, anapproach is described to determine resistance under similar conditions,namely in assisted ventilation. This prior approach consists of applyingat least two different types of transient changes in flow in the courseof the inflation phase of the ventilator. The changes in airway pressure(Paw), flow ({dot over (V)}), and volume (V) during these transient flowchanges are compared with the time course of these variables inunperturbed breaths. While this approach is capable of providingaccurate information about R, it has several limitations. First, becauseof considerable breath-by-breath variability in the time course of Paw,{dot over (V)} and V in spontaneous unperturbed breaths, it is necessaryto average large numbers of perturbed and unperturbed breaths in orderto arrive at the real change that occurred during the perturbation.Accordingly, information about resistance is delayed until asufficiently large number of observations has been averaged.Furthermore, for the same reasons, any true change in patient'sresistance is not detected in a timely way. Second, this approachrequires at least two different kinds of perturbations. Because, asindicated earlier, a large number of observations is required with eachperturbation, this requirement delays the acquisition of reliableinformation further. Third, the need to average large numbers of breathsand a large number of data points from each breath, greatly increasesthe computing and storage requirements of the computer used to processthe information to provide the value of R. This requirement adds furtherstrain on the extensive and highly complex operations carried out bymodern, computer controlled ventilators.

SUMMARY OF INVENTION

The method and apparatus described in detail herein in accordance withthe present invention, represent a considerable simplification of theapproach proposed by Younes in U.S. Pat. No. 5,884,622. As indicatedabove, the main obstacle to determining respiratory resistance duringassisted ventilation is the uncertainty about what happens to Pmusduring interventions in the course of the inflation phase of theventilator. The comparison between perturbed and unperturbed breaths wasthe approach used in U.S. Pat. No. 5,884,622. By contrast, in accordancewith the present invention, the behavior of Pmus during the interventionis predicted from estimates of the change in Pmus in the intervalimmediately preceding the intervention. In this manner, all the requiredinformation necessary to determine R can be obtained from a singleintervention in a single breath. This approach greatly reduces thecomputational requirements necessary to determine R, and the timerequired to obtain information that is clinically useful, such as inassisted ventilation

In accordance with the present invention, respiratory resistance (R) isdetermined while allowing for the presence of pressure generated byrespiratory muscles (Pmus) but without requiring knowledge of its actualvalue or an accurate value of passive respiratory elastance (E).

In accordance with one aspect of the present invention, there isprovided a method of determining respiratory system resistance (R) in apatient receiving gas from a ventilatory assist device (ventilator),comprising estimating the flow rate ({dot over (V)}) and volume (V) ofgas received by the patient from the ventilator, estimating pressurenear the airway of the patient (Paw), generating a signal that resultsin a step decrease (negative pulse) in the pressure and/or flow outputof the ventilator during selected inflation cycles, measuring Paw, {dotover (V)} and V at a point (T_(O)) near the beginning of the pulse(Paw_(O), {dot over (V)}, V_(O)), at a point (T₁) near the trough of thenegative pulse (Paw_(I), {dot over (V)}_(I), V_(I)), and at a point(T⁻¹) preceding T_(O) but after the onset of inspiratory effort (Paw⁻¹,{dot over (V)}⁻¹, V⁻¹); and calculating the value of resistance (R) fromthe differences between Paw, {dot over (V)} and V at T_(O) and at T_(I)and where the change in patient generated pressure (Pmus) in theinterval T_(O)→T_(I)(ΔPmus(T_(O)→T_(I))) is estimated by extrapolation,from the differences between Paw, {dot over (V)} and V at T_(O) and atT⁻¹, in accordance with equation (8).

As described in more detail below, the present invention includesmodifications to the method as alternative steps to determining R.

In accordance with another aspect of the present invention, there isprovided an apparatus which interfaces with ventilatory assist devices(ventilators) determining respiratory system resistance (R), comprisinga flowmeter, with associated electronic circuitry, that estimates theflow rate ({dot over (V)}) and volume (V) of gas received by a patient,a pressure sensor that estimates pressure near the airway of the patient(Paw), and electronic circuitry which receives the Paw, {dot over (V)}and V signals from above mentioned circuitry and which is also connectedto the control system of the ventilator, comprising:

-   -   circuitry that generates an output that results in a step        decrease (negative pulse) in the pressure and/or flow output of        the ventilator during selected inflation cycles;    -   circuitry that measures Paw, {dot over (V)} and V at a point        (T_(O)) near the beginning of the pulse (Paw_(O), {dot over        (V)}, V_(O)), at a point (T₁) near the trough of the negative        pulse (Paw₁, {dot over (V)}₁, V₁), and at a point (T⁻¹)        preceding T_(O) but after the onset of inspiratory effort        (Paw⁻¹, {dot over (V)}⁻¹, V⁻¹);    -   circuitry to calculate the value of resistance (R) from the        differences between Paw, {dot over (V)} and V at T_(O) and at        T_(I) and where the change in patient generated pressure (Pmus)        in the interval T_(O)→T_(I)(ΔPmus(T_(O)→T₁)) is estimated, by        extrapolation, from the differences between Paw, {dot over (V)}        and V at T_(O) and at T⁻¹, in accordance with equation (8).

As described in more detail below, the present invention includesmodifications to the apparatus as alternative combinations of elementsto determine R.

BRIEF DESCRIPTION OF DRAWINGS

FIG. 1 shows a tracing of airway presence (Paw), flow and volume showinga negative pulse and the three times at which measurements are taken;

FIG. 2 is a schematic representation of apparatus for carrying out themethod in accordance with a preferred embodiment of the invention; and

FIG. 3 shows schematically the various elements of a micro controllerused in connection with the apparatus of FIG. 2.

GENERAL DESCRIPTION OF THE INVENTION

According to the equation of motion, the total pressure applied to therespiratory system (P_(appl)) is dissipated against elastic, resistiveand inertial opposing forces. Thus:P _(appl) =P _(el) +P _(res) +P _(iner)where:

-   -   P_(el) is elastic recoil pressure and is given by the product of        volume above passive functional residual capacity (FRC) (V) and        elastance (E); P_(el)=V·E,    -   P_(res) is the pressure dissipated against resistance and is        given by the product of flow ({dot over (V)}) and R;        P_(res)={dot over (V)}·R, and,

P_(iner) is the pressure dissipated against inertia and is given by theproduct of flow accelaration (the rate of change in flow in I/sec²;{umlaut over (V)}) and inertia (I). Because I of the respiratory systemis very small (≈0.02 cmH₂O/l/sec²), P_(iner) can be ignored so long asmeasurements are made at relatively low {umlaut over (V)} (e.g. <10I/sec²). In mechanically ventilated patients, {umlaut over (V)} mayexceed this level only in the first about 100 to 200 msec of theinflation phase during volume cycled and high level pressure supportventilation (PSV). Accordingly, by avoiding measurements in this region,the equation solved to calculate P can be simplified by neglectingP_(iner).

During assisted ventilation, P_(appl) is made up of two components, oneprovided by the ventilator (Paw) and one provided by the patient (Pmus).Thus, P_(appl)=Paw+Pmus. With this equation and earlier considerations,the equation of motion can be rewritten and rearranged as follows asequation (1):{dot over (V)}·R=Paw+Pmus−V·E  (1)To the extent that Pmus at a given instant is not known, accurateelastance values may not be available and V, relative to passive (FRC),is also not known (in view of possible dynamic hyperinflation or activereduction in volume below FRC by expiratory muscles), it is not possibleto solve for R using a set of measurements made at one point during theinflation phase. For this reason, any approach to measure resistanceduring inflation in such patients must involve measurements at more thanone point, having different flow values, as described herein.

In one aspect of the present invention, Paw (and hence flow) is rapidlyreduced (negative pulse) during the inflation phase (FIG. 1). Primarymeasurements of Paw, {dot over (V)} and V are made at or near the pointwhere Paw and flow begin declining (T_(O)) and at or near the trough ofpressure during the negative pulse (T_(I)). These two sampling pointsare chosen because ΔP/Δt and ΔV/Δt are minimal. In this fashion,inertial forces can continue to be ignored. More importantly, errorsrelated to differences in delay and frequency response of the pressureand flow measuring systems can be avoided. This advantage isparticularly relevant since, in modern ventilators, Paw and patient floware not measured directly near the ET tube but are estimated from remotesites and, hence, the signals may be subject to different delays andresponse characteristics. Even minor differences in these properties cancause serious errors when Paw and patient flow are changing rapidly (forexample, during the declining phase of the pulse).

Equation 1 can be written for T_(O) and T_(I) as follows, as equations(2) and (3):{dot over (V)} _((O)) ·R=Paw _((O)) +Pmus _((O)) −V ₍₀₎ ·E  (2){dot over (V)} ₍₁₎ ·R=Paw ₍₁₎ +Pmus ₍₁₎ −V ₍₁₎ ·E  (3)Subtracting equation 3 from equation 2 yields equation (4):R({dot over (V)} _((O)) −{dot over (V)} ₍₁₎)=(Paw _((O)) −Paw ₍₁₎)+(Pmus_((O)) −Pmus ₍₁₎)−E(V _((O)) −V ₍₁₎)  (4)Rearranging equation (4) to solve for R:R=[(Paw _((O)) −Paw ₍₁₎)+ΔPmus(T _(O) →T _(I))−E(V _((O)) −V ₍₁₎)]/({dotover (V)} _((O)) −{dot over (V)} ₍₁₎)In theory, if the time interval between T_(O) and T_(I) (i.e. Δt) isinfinitely small, the differences in Pmus and in volume can be ignoredand ΔPaw becomes ΔPres. In practice, however, during mechanicalventilation it is not possible to instantly reduce flow from one valueto another relatively stable value (i.e. at which ΔP/Δt and Δ{dot over(V)}/Δt are acceptably small). Even if flow exiting the ventilator isaltered suddenly, a finite time must elapse before the flow to thepatient stabilizes at the new value in view of continued flow from thetubing to patient in the process of decompression of the circuit. Δt,therefore, cannot be made short enough to ignore changes in Pmus betweenT_(O) and T₁ and ΔPmus in this interval has to be accounted for.

In this aspect of the present invention, the change in Pmus betweenT_(O) and T₁ is estimated by assuming that Pmus changes in this timerange at the same rate as in the period immediately preceding T_(O).This is not an unreasonable assumption if the time interval betweenT_(O) and T₁ is relatively brief (for example, approximately 100 msec).The rate of change in Pmus immediately before the pulse is estimated bysampling Paw, {dot over (V)} and V at a point shortly before T₀ (forexample, 100 msec prior to T_(O))(T_(−I)), (FIG. 1). The following twoequations (5) and (6) provide estimates of Pmus at T_(O) and T⁻¹respectively and represent rearrangement of the equation of motion(equation (1)):Pmus _(O) =V _(O) ·E+{dot over (V)} _(O) ·R−Paw _(O)  (5)Pmus _(−I) =V _(−I) ·E+{dot over (V)} ⁻¹ ·R−Paw ⁻¹  (6)

Subtracting equation 6 from equation 5 and dividing by Δt−1 (timebetween T_(O) and T⁻¹) gives ΔPmus/Δt in the interval (Δt₁) prior toT_(O) according to equation (7):ΔPmus/Δt=(1/Δt _(−I))[E(V _(O) −V _(−I))+R({dot over (V)} _(O) −{dotover (V)} ⁻¹)−Paw _(O) +Paw _(−I)]  (7).

Assuming that Pmus changes at the same rate between T_(O) and T_(I), thechange in Pmus between these two points is given by:(Pmus _(O) −Pmus ₁)=−Δt _(I)  [equation 7]where Δt_(I) is the time interval between T₁ and T_(O).

Substituting [−Δt (equation 7)] for (Pmus_(O)−Pmus₁) in equation 4 andrearranging provides equation (8):R=[(Paw _(O) −Paw _(I)+(Δt/Δt ⁻¹)(Paw _(O) −Paw ⁻¹))−E(V _(O) −V₁+(Δt/Δt ⁻¹)(V _(O) −V ⁻¹))]/({dot over (V)} _(O) −{dot over (V)}₁+(Δt/Δt ⁻¹)({dot over (V)} _(O) −{dot over (V)} ⁻¹))  (8)

The only unknown in the numerator of equation 8, which is the estimateof Pres, is E. However, unlike the case in equation 4, the difference inV between T_(O) and T₁ is now reduced by the term (Δt₁/Δt⁻¹)(V_(O)−V⁻¹).If the pulse is initiated during the rising phase of flow (e.g. FIG. 1),average {dot over (V)} in the intervals T⁻¹ to T_(O) and T_(O) to T₁will not be substantially different and, given that the time intervalsbetween T_(O) and T_(I) and between T_(O) and T₁ are quite small (ca 0.1sec), the entire volume term is reduced to nearly zero. Under theseconditions, any errors in estimating E should result in very minorerrors in estimating Pres, and hence resistance (R), and, in the absenceof a known value of E, a default value, representing, for example,average E in ventilator dependent patients, can be used without muchrisk of significant errors. It should also be noted that, because allvolume points are obtained from the same breath, differences between anytwo volume values represent differences in absolute volume, relative topassive FRC. As a result, offsets of volume, relative to passive FRC, atthe beginning of the breath become irrelevant.

The above derivation of equation 8 entails the assumption that the valueof R is constant or, specifically, that R is independent of flow andvolume. In reality, R may vary with flow, particularly in intubatedpatients, if only because the resistance of the endotracheal tubeincreases with flow. Likewise, R may be dependent on lung volume in somepatients. Equation 8 can be adapted to allow for R being flow and/orvolume dependent. The number of mathematical functions that can be usedto characterize flow or volume dependence of R is infinite. It would beimpractical to provide formulations of equation 8 that allow for allconceivable mathematical descriptions of flow and/or volume dependence.Rather, one example will be illustrated which represents the most widelyaccepted behavior of R in mechanically ventilated intubated patients,namely that R is minimally (or not at all) affected by volume but thatit increases with flow according to Rohrer's equation (R=K₁+K₂{dot over(V)}). It is recognized, however, that any individual with modestmathematical skills can utilize the same information obtained in thisaspect of the present invention (i.e. Paw, V and {dot over (V)},measured at T_(O), T_(I) and at points preceding T_(O)) to derive thepressure-flow relation where mathematical functions other than Rohrer'sequation are assumed to apply.

The following equations (2a to 8a) correspond to equations 2 to 8 aboveafter making appropriate modifications to allow for R to increase withflow according to Rohrer's equation (R=K₁+K₂{dot over (V)}):

 K ₁ {dot over (V)} _(O) +K ₂ {dot over (V)} _(O) ² =Paw _(O) +Pmus _(O)−V _(O) ·E  (2a)K ₁ {dot over (V)} ₁ +K ₂ {dot over (V)} _(O) ² =Paw ₁ +Pmus ₁ −V ₁·E  (3a)K ₁({dot over (V)} _(O) −{dot over (V)} _(I))+K ₂({dot over (V)} _(O) ²−{dot over (V)} ₁ ²)=(Paw _(O) −Paw ₁)+(Pmus _(O) −Pmus _(I))−E(V _(O)−V ₁)  (4a)Pmus _(O) =V _(O) ·E+{dot over (V)} _(O) ·K ₁ +V _(O) ² ·K ₂ −Paw_(O)  (5a)Pmus _(I) =V _(−I) ·E+{dot over (V)} _(−I) ·K ₁ +V ⁻¹ ² ·K ₂ −Paw⁻¹  (6a)ΔPmus/Δt=(1/Δt ⁻¹)[E(V _(O) −V ₁)+K ₁({dot over (V)} _(O) −{dot over(V)} _(−I))+K ₂({dot over (V)} _(O) ² −{dot over (V)} _(−I) ²)−Paw _(O)+Paw ⁻¹]  (7a)K ₁({dot over (V)} _(O) −{dot over (V)} ₁+(Δt ₁ /Δt ⁻¹)({dot over(V)}_(O) −{dot over (V)}⁻¹))+K ₂({dot over (V)} _(O) ² −{dot over (V)} _(I) ²+(Δt 1 /Δt−1)({dotover (V)}_(O) ² −{dot over (V)} _(O−I) ²)=(Paw _(O)−Paw_(I)+(Δt/Δt ⁻¹)(Paw _(O) −Paw ⁻¹))−E(V _(O) −V _(I)+(Δt/Δ⁻¹)(V _(O) −V ⁻¹))  (8a)

From each applied pulse, an equation of the form of equation (9)accordingly results:K ₁ ·X+K ₂ ·Y=Z  (9)where X is the flow term (first bracketed term to left of equation 8a),Y is the {dot over (V)}² term (second bracketed term in equation 8) andZ is the Pres term (right side of equation 8). To obtain Z, a knownvalue of E is used or, in the absence of this information, a defaultvalue (e.g. 28 cmH₂O/l, representing average E in mechanicallyventilated patients (personal observations), may be used. Resistance canbe obtained from the above equation (9) in one of several ways. Some ofthese are listed below:

1) K₂ is initially assumed to be zero and resistance is estimated fromZ/X. The resistance value obtained in this fashion represents the slopeof the P {dot over (V)} relation between {dot over (V)}_(O) and weightedaverage of {dot over (V)}_(I) and {dot over (V)}_(−I). If {dot over(V)}_(I) and {dot over (V)}⁻¹ are not substantially different (e.g. FIG.1), R calculated in this fashion can be assumed to represent the slopeof the P {dot over (V)} relation between {dot over (V)}_(O) and either{dot over (V)}₁ or {dot over (V)}_(−I) or the mathematical average ofthe two. It can be shown, using Rohrer's equation, that the slope of theP {dot over (V)} relation between any two flow points (incrementalresistance, IR) is the same as the resistance at a flow corresponding tothe sum (flow-sum) of the two flow points (in this case ({dot over(V)}_(O)+{dot over (V)}₁)). With this treatment, R is reported asresistance at a specific flow (i.e. flow sum).

2) If a range of flow-sum values is obtained in successive pulses,either spontaneously or by design, a range of IR values will alsoresult. A regression between IR (dependent variable) and flow-sum willresult in a significant correlation if a sufficiently wide range offlow-sum is present. The intercept of this regression is K_(I) and theslope is K₂. These can be reported as such. Alternatively, resistancemay be reported as the sum of K_(I) and K₂, which is resistance at astandard flow of 1.0 l/sec. This has the advantage that changes inreported resistance reflect real changes in resistance whereas withapproach #1, alone, the reported resistance may change simply becauseflow is different.

3) The values of K₁ and K₂ can be derived from the results of two pulseshaving different X and Y values, or by regression analysis of theresults of multiple pulses displaying a range of X and Y values. Theprocedure of applying pulses can be deliberately planned to result in awide range of X and Y values in order to facilitate this analysis. Forexample, pulses may be initiated at different flow rates, so that {dotover (V)}_(O) is variable, and/or the decrease in {dot over (V)} duringthe pulse can also be deliberately varied, to result in a range of {dotover (V)}₁.

4) In the absence of reliable, directly determined K₁ and K₂ values,following approach #2 above, K₂ can be assumed to equal K₂ of theendotracheal tube (ET) and equation 9 is solved for K₁. Thus,K₁=(Z−(Y·K₂ET))/X. The K₂ values of clean ET tubes of different sizesare widely available. Resistance can be reported as K₁+K₂ET, reflectingresistance at a standard flow of 1.0 l/sec. The resistance so reportedmay differ from actual resistance at 1 l/sec to the extent that actualK₂ET may differ from the assumed K₂ of a clean tube AND the flow atwhich R estimates are made are different from 1.0 l/sec. The error inestimated resistance (at 1 l/sec), if actual K₂ (K₂ actual) is differentfrom assumed K₂ is given by R_(error)=(K₂ actual−K₂ assumed)(1−Y/X). Itcan be seen that the error in estimating R at 1 l/sec using an assumedK₂ is a fraction of the difference between the actual and assumed K₂value.

Potential Sources of Errors and Approaches to Minimize Such Errors:

1) Measurement Noise:

In mechanically ventilated patients, the Paw and {dot over (V)} signalsare subject to noise from multiple sources. These include airwaysecretions, cardiac artifacts, liquid in the tubing and oscillations orvibrations in the flow delivery system of the ventilator. The noise inthe Paw signal may be in phase or out of phase with that in the {dotover (V)} signal depending on the source of noise and the frequencyresponse of the two measuring systems. Out of phase noise has a greaterimpact on estimated R particularly if the critical measurements (e.g. atT_(O), T_(I) and T⁻¹) are obtained from discrete points of unfilteredsignals. Such noise results in an increased random variability ofestimated R in successive measurements. A more systematic error mayresult if the pulse is programmed to begin when a certain flow isreached. Here, there is an increased probability that the pulse willbegin on the upswing of a positive flow artifact.

Errors related to measurement noise can be reduced by a variety ofapproaches:

a) The most effective approach is to insure that the change in flowproduced by the intervention (i.e. change in flow between T_(O) andT_(I)) is large relative to the amplitude of the noise.

b) Elimination of sources of noise to the extent possible.

c) Critical filtering of the Paw and {dot over (V)} signals.

d) To minimize systematic errors, the pulse should preferably not beginwhen a fixed level of flow is reached (see above).

e) Averaging the resistance results obtained from a number of pulses.

2) Difference in Response Characteristics of Paw and {dot over (V)}Measuring Systems:

Difference in response characteristics of the measuring systems causesthe peak and trough of the measured pressure to occur at different timesrelative to the flow signal even if the peaks and troughs of the twosignals were, in reality, simultaneous. If T_(O) is taken as the time ofpeak Paw, flow at T_(O) will underestimate real flow, and vice versaAlso, such differences convert the relatively innocuous in-phaseoscillations originating from ventilator flow delivery systems topotentially more serious out-of-phase oscillations in Paw and flow. Tominimize the impact of these differences, the phase lag between the Pawand flow measuring systems should be as short as possible over thefrequencies of interest. In addition, the pulse can be designed to avoidsharp peaks and troughs.

3) Errors Related to Extrapolation of the Pmus Trajectory:

These are potentially the most serious particularly when respiratorydrive, and hence ΔPmus/Δt, is high. The proposed approach involves theassumption that ΔPmus/Δt during the pulse is the same as ΔPmus/Δt over afinite period prior to the pulse. This assumption can be in error for avariety of reasons. These, and possible ways to minimize these potentialerrors, are discussed below:

a) Termination of inspiratory effort (neural T_(i)) during the pulse:This can potentially produce the largest errors in estimated R. Thus,assume that ΔPmus/Δt prior to T_(O) is 40 cmH₂O/sec and Δt_(I) (i.e.T_(I)−T_(O)) is 0.15 sec. The estimated increase in Pmus between T_(O)and T₁ would be 6 cmH₂O. If, however, neural T_(i) ends near T_(O), Pmuswill decrease instead of increasing. Because the rate of decline in Pmusduring neural expiration is fastest soon after the end of neural T_(i),the decrease in Pmus may actually be greater than the assumedextrapolated increase, with the error in estimated ΔPmus being >12cmH₂O. It can be seen from equation 4 that this condition translatesinto an error of corresponding magnitude in estimated Pres. If thedifference between {dot over (V)}_(O) and {dot over (V)}₁ is 0.4 l/sec,this error would translate into an error of >30 cmH₂O/l/sec in estimatedresistance.

Because of the potentially large magnitude of this error, it isnecessary to insure that peak Pmus (end of T_(i)) does not occur betweenT_(O) and T_(I). This condition is easy to accomplish duringProportional Assist Ventilation (PAV). In this mode, the end ofventilator cycle is automatically synchronized with patient effort andis constrained to occur during the declining phase of Pmus. So long aspulses are not delivered in the last fraction (ca 30%) of ventilator T₁,one is assured that T_(i) termination did not occur within the pulse.With pressure support ventilation (PSV) and assisted volume-cycledventilation, such synchrony is not assured, however, and T_(i) mayterminate at any point within or even beyond the inflation phase. T_(i)termination may occur, per chance, during some of the pulses resultingin errors of differing magnitudes depending on ΔPmus/Δt prior to thepulse, the point within the pulse at which T_(i) terminated, the rate ofdecline in Pmus beyond the peak, and the difference in flow betweenT_(O) and T₁. Considerable variability may occur between the results ofdifferent pulses. For this reason, application of this approach duringPSV and volume cycled ventilation may produce less reliable resistancevalues.

b) Shape of the rising phase of Pmus: The rate of rise of Pmus duringthe rising phase is not constant. Differences between ΔPmus/Δt in theinterval T_(O) to T_(I) (i.e. Δt₁) and T_(−I) to T_(O) (i.e. Δt⁻¹)causes errors in estimated R for the same reasons discussed under (a)above. A Δt of approximately 0.1 sec is both feasible and consistentwith minimal errors related to response characteristics of the measuringsystems. It is unlikely that an important change in ΔPmus/Δt would occurover this brief time interval, provided all measurement points (i.e.T_(O), T_(I), T_(−I)) occur during either the rising or declining phaseof Pmus. What needs to be avoided is the occurrence of T_(i) terminationbetween T_(−I) and T_(I) and this can be accomplished by insuring thatpulse application occurs either early in the inflation phase or verylate in the inflation phase in the PAV mode. In this mode, there isassurance that pulses applied in the first 50% of the inflation phaseoccur, in totality (i.e. T⁻¹ to T₁), on the rising phase of Pmus whilepulses applied very near the end of the inflation phase will occur intotality on the declining phase of Pmus. In either case, there is littlelikelihood of a major change in ΔPmus/Δt over the brief period of thepulse and extrapolation from one segment to the next, within the briefpulse period, should not result in significant errors.

c) Behavioral responses: The change in Pmus following the initiation ofthe pulse may deviate dramatically from that expected from the precedingtime interval if the patient perceives the pulse and reacts behaviorallyto it. The minimum latency for behavioral responses to changes in Pawand flow is approximately 0.2 sec, even in very alert normal subjects.It follows that errors related to perception of the pulse, withconsequent behavioral responses, can be avoided if measurements arerestricted to the approximately 0.2 sec interval after initiation of thepulse. Behavioral responses, however, can occur without perception ifthe change is anticipated. For example, if a perturbation occursregularly every 5 breaths, the patient may alter his/her respiratoryoutput every fifth breath, even before the pulse is initiated. Theoccurrence of anticipatory responses can be minimized by randomizing theorder of pulse applications.

d) Non-behavioral neuromuscular responses to changes in flow: The rapidreduction in flow in the course of an ongoing inspiratory effort may,theoretically, elicit reflex changes in neural output with much shorterlatencies than behavioral responses. In addition, the change in flowand, consequently, in time course of volume, may elicit changes in Pmus,independent of changes in electrical activation, through the operationof the intrinsic properties of respiratory muscles (force-length andforce-velocity relations). An important contribution from either ofthese responses following the onset of the pulse (between T_(O) andT_(I)) could alter the time course of Pmus relative to the coursepredicted from the pre-pulse interval and introduce errors in estimatedPres. Based on experimental results, these effects are likely to besmall if the change in flow is modest (for example, <1 l/sec) and theintervention is carried out early in the inflation phase where Pmus isrelatively low.

e) Pmus noise: Noise in the Pmus signal can introduce errors when thechange in Pmus over a relatively brief period (for example, 0.1 sec) isused to estimate the change in a subsequent interval. Pmus noise can bereal or artifactual. Tracings of P_(di) (transdiaphragmatic pressure),for example, often have a jagged rising phase. Furthermore, when Pmus isestimated from P, {dot over (V)} and V, as opposed to being measured,independent noise in the pressure and flow signals (for example, cardiacartifact, secretions . . . etc) can introduce noise in estimated Pmus,even if the true rising phase of Pmus is smooth. The impact of Pmusnoise on estimated resistance is the same whether the noise is real orartifactual. Random noise in the Pmus signal may be expected to increasevariability in measured resistance values, reducing the reliability ofinformation obtained from single pulses. This condition can be dealtwith by averaging the results of several observations over a number ofbreaths. Furthermore, the impact of Pmus noise can also be reduced byusing a relatively large change in flow between T_(O) and T₁.

Alternative Approaches to Calculation of Resistance Using the PulseTechnique:

1) Estimation of the change in Pmus during the pulse using aninterpolation approach:

In the above description, the change in Pmus between T_(O) and T₁ (i.e.ΔPmus(T_(O)→T_(I))) was estimated by extrapolation of the Pmustrajectory in the interval T⁻¹ to T_(O). An alternative approach is toestimate ΔPmus(T_(O)→T₁) by interpolation between two points, one before(for example, at T_(O)) and one after the trough of Paw (T₂). In thiscase, Paw, {dot over (V)} and V are measured at T_(O) (i.e. Paw_(O),{dot over (V)}_(O) and V_(O)) and at T₂ (i.e. Paw₂, {dot over (V)}₂ andV₂) in addition to at T₁. T₂ should preferably be chosen at a point,after T₁, where ΔPaw/Δt and/or Δ{dot over (V)}/Δt are very small tominimize inertial forces. With this alternative approach, equation 8 canbe written as follows:R=[(Paw _(O) −Paw ₁+(Δt ₁ /Δt ₂)(Paw ₂ −Paw _(O)))−E(V _(O) −V _(I)+(Δt_(I) /Δt ₂)(V ₂ −V _(O)))]/({dot over (V)}_(O) −{dot over (V)} ₁+(Δt_(I) /Δt ₂)({dot over (V)}₂ −{dot over (V)} _(O)))  (8 inter)where Δt₂ is the interval between T₂ and T_(O). Equation 8a can bewritten as follows:K ₁({dot over (V)} _(O) −{dot over (V)} ₁+((Δt ₁ /Δt ₂)({dot over (V)} ₂−{dot over (V)} _(O))))+K ₂({dot over (V)} _(O) ² −{dot over (V)} ₁²+((Δt ₁ /Δ ₂)({dot over (V)} ₂ ² −{dot over (V)} _(O) ²)))=(Paw _(O) −Paw ₁+((Δt ₁ /Δt ₂)(Paw ₂ −Paw _(O))))−E(V _(O) −V ₁+((Δt_(I) /Δt ₂)(V ₂ −V _(O))))  (8a inter)There are advantages and disadvantages to the interpolation approach,relative to the extrapolation approach described earlier. The mainadvantage is that, in principle, estimating an unknown value byinterpolation between values before and after the unknown value is moreaccurate than estimating the unknown value through extrapolation of datapoints which are all occurring before or after the unknown value. Thepractical disadvantages in this particular application, however, arethat point T₂ occurs beyond the pulse intervention and, as well, laterin inspiration. Pmus at T₂ may thus be corrupted through behavioral orreflex responses to the preceding intervention, and by the greaterlikelihood that termination of inspiratory effort with precipitousdecrease in Pmus, may occur prior to T₂,

2) Combined use of the extrapolation and interpolation techniques:

R can be estimated using both the extrapolation technique (equation 8 or8a) and the interpolation technique (equation 8 inter and 8a inter) andthe results of the two approaches may be averaged using a suitableaveraging technique.

While either the interpolation approach or the combined approach may beused in preference to the extrapolation technique, my practicalexperience favors the extrapolation technique. Thus, it was found instudies on 67 ventilator dependent patients that the results of theextrapolation approach are in closer agreement with results obtainedduring controlled ventilation than the results of the other twoalternative approaches.

3) Use of back extrapolation, instead of forward extrapolation, of Pmus:

The change in Pmus between T_(O) and T₁ can be estimated by backextrapolation of data from a period following T₁. Thus, Paw, {dot over(V)} and V are measured at T₂ (see alternative approach #1 above).Equation 8 and 8a are modified to reflect these sampling points asfollows:R=[(Paw _(O) −Paw ₁+(Δt _(I) /Δt ₂)(Paw ₂ −Paw ₁))−E(V _(O) −V ₁+(Δt_(I) /Δt ₂)(V ₂ −V _(I)))]/({dot over (V)} _(O) −{dot over (V)} ₁+(Δt ₁/Δt ₂)({dot over (V)} ₂ −{dot over (V)} ₁))   equation 8 (bextra)andK ₁({dot over (V)} _(O) −{dot over (V)} _(I)+(Δt ₁ /Δt ₂)({dot over (V)}₂ −{dot over (V)} _(I))))+K ₂({dot over (V)} _(O) ² −{dot over (V)} _(I)²+((Δt _(I) /Δt ₂)({dot over (V)} ₂ ² −{dot over (V)} ₁ ²)))=(Paw _(O) −Paw ₁+((Δt ₁ /Δt ₂)(Paw ₂ −Paw ₁)))−E(V _(O) −V ₁+((Δt _(I)/Δt ₂)(V ₂ −V _(I))))  equation 8a (bextra)

4) ΔPmus/Δt prior to T_(O) can be estimated from values obtained at twopoints

both of which occurring before T_(O) (e.g. T⁻¹ and T⁻²). Althoughfeasible and should provide reasonably accurate results, it has littleadvantage over the use of T_(O) and only one preceding point, whileadding more computation complexities.

5) Use of regression analysis to estimate ΔPmus/Δt prior T_(O) or beyondT₁: The extrapolation approach described above utilizes measurements atonly two points (e.g. T_(O) and T⁻¹) to estimate ΔPmus/Δt(T_(O)→T_(I)).Although computationally much more intensive, ΔPmus/Δt prior to theonset of the pulse, or between T₁ and T₂, can be estimated by samplingPaw, {dot over (V)} and V at multiple points prior to, or after, thepulse and estimating ΔPmus/Δt by suitable regression analysis. Thestandard equations for linear and non-linear regression can be appliedto the multiple data sets, to obtain an estimate of Paw, {dot over (V)}and V at T₁. These are then inserted at the appropriate locations inequations 8 and 8a.

6) Use of positive flow pulses (transients): Whereas there is describedabove the application of the procedure of the invention using negativePaw and flow transients (for example, FIG. 1), the same approach can beapplied to imposed positive flow and Paw transients. Here, Paw, {dotover (V)} and volume are also measured immediately before theperturbation (T_(O)), at or near the point of maximum Paw (or flow) ofthe positive pulse (T_(I)) (as opposed to the trough of the negativepulse) and at a third point, either before T_(O), to implement theextrapolation technique, or after T₁, to implement the interpolation orback extrapolation techniques. The values of Paw, {dot over (V)} and Vobtained at the three points (T_(O), T₁ and T⁻¹ or T_(O), T₁ and T₂) arethen inserted in equation 8 or 8a (for extrapolation approach), 8 interor 8a inter (for interpolation approach) or 8 (bextra) and 8a (bextra)(for the back extrapolation approach). Regression analysis can also beused on multiple data prior to T_(O). In my experience, negative pulsesprovide more reliable results and are, therefore, preferred. The morereliable result using negative pulse is likely related to two factors.First, negative pulses dictate the occurrence of a point at whichΔPaw/Δt is zero, which can be used as T_(O) (see FIG. 1). With positivepulses, this cannot be assured. There are advantages to making themeasurements at points where ≢Paw/Δt and Δ{dot over (V)}/Δt are nearzero, (as discussed above). Second, in many patients there aresubstantial differences in time of occurrence of peak flow and peak Pawwhen positive pulses are given, which introduces uncertainty in theresults.

DESCRIPTION OF PREFERRED EMBODIMENT

FIG. 2 shows an overview of a preferred embodiment of apparatus forcarrying out the present invention. This preferred embodiment reflectsthe actual system used to validate the inventive procedures of theinvention in 67 ventilator-dependent patients. The preferred embodimenthas several components. Although in FIG. 2, these components are shownas distinct from each other, such representation is for the sake ofillustration of these components, in actual practice all threecomponents can be incorporated within a single unit (the ventilator).

A gas delivery unit 10 is a ventilator system that is capable ofdelivering proportional assist ventilation (PAV). A variety ofmechanical systems can be used to deliver PAV and some are commerciallyavailable, which use blower-based, piston-based and proportionalsolenoid systems. PAV is described in U.S. Pat. No. 5,107,830 (Younes),assigned to the assignee hereof and the disclosure of which isincorprorated herein by reference. The ventilator illustrated in thepreferred embodiment consists of a piston 12 reciprocating within achamber 14. The piston 12 is coupled to a motor 16 that applies force tothe piston 12 in proportion to input received from the PAV pressurecontrol unit 18. A potentiometer 20 measures the piston displacementwhich corresponds to the volume change during the ventilator cycle.After certain corrections related to leaks and gas compression, thissignal conveys the amount of gas (volume) received by the patient. Thevolume signal (V) is differentiated using a suitable differentiator 22to result in a flow signal ({dot over (V)}). The PAV pressure controlunit 18 generates a signal that is the sum of a suitably amplified flowsignal and a suitably amplified volume signal with amplification factorsbeing set by the user, which signal is used to control the motor 16. Thepiston chamber 14 receives suitable gas mixture through an inlet port 24and delivers gas to the patient through an outlet port. Duringinspiration, an exhalation valve control circuit closes the exhalationvalve 26 ensuring that the gas pumped by the piston 12 is delivered tothe patient through valve 27. At the end of the inspiratory cycle, theexhalation valve control circuit opens the exhalation valve 26 to allowexpiratory flow to occur prior to the next cycle.

A micro controller 28 receives the flow and airway pressure signals.These can be obtained directly from ventilator outputs of flow ({dotover (V)}) out and airway pressure (P). Alternatively, flow and airwaypressure are measured independently by inserting a pneumotachograph (30)and an airway pressure outlet between the Y connector 32 and thepatient. The latter approach is the one illustrated in FIG. 2. Pressuretransducers are provided (FT and PT) to generate signals in proportionto airflow and airway pressure near the endotracheal tube 34. Althoughthis is a more direct way of estimating patient flow and airwaypressure, reasonably accurate estimates can be obtained from sensorswithin the ventilator body, remote from the patient, after allowancesare made for tube compression. The patient flow and airway pressuresignals are continuously monitored by the micro controller 28. At randomintervals, electric pulses are generated by the micro controller and areconveyed to the PAV delivery system via suitable output ports. Thesepulses may be either negative or positive, as described above. The pulseoutput is connected to the PAV pressure control unit 18 within the PAVdelivery system via 36. The electrical pulse results in a temporarydecrease or increase in the output of the PAV pressure control unitrelative to the output dictated by the PAV algorithm. This, then,results in either a corresponding decrease or increase in airwaypressure for a brief period (for example, approximately 0.2 sec), at atime determined by the micro controller, in selected breaths.

The basic components of the micro controller 28 used in this preferredembodiment are shown in FIG. 3. The flow and airway pressure signals arepassed to signal conditioning circuits ((LM324OP-AMPS) or equivalent)which condition the input voltage signals into 0 to +5 volts. The twosignals are passed through to an analog to digital convertor on themicro controller. The digitized flow signal is integrated to provideinspired volume. A clock circuit allows flow, pressure and volume to besampled at precise intervals. The basic computer is an MC68HC16 withAM29F010 ROM and KM68-1000 RAM. A preferred embodiment of the mastercomputer program includes several functions as follows:

-   (1) A function to identify the beginning of inspiration. Inspiration    is deemed to have started when inspiratory flow exceeds a certain    threshold (e.g. 0.1 l/sec) and remains above this level for a period    of at least about 150 msec beyond this point.-   (2) A random number generator function generates a number between 4    and 11 which determines the number of breaths between any two    successive perturbations. This results in an average of 6    unperturbed breaths between any two successive perturbations. Any    other convenient integers and average may be chosen. The average    number also can be over-ridden by the user through a manual input    via a key pad. The user may elect to deliver the perturbations at a    faster average rate to speed up the data collection or, conversely,    the frequency of application of perturbations can be slowed down,    as, for example, when the clinical condition is fairly stable.    Clearly other methods of ensuring that pulses are applied at random    intervals are possible. Pulses may also be applied at regular    intervals, although this may result in anticipatory responses by the    patient which may corrupt the measurements under some circumstances.-   (3) An event processor function which controls the time of    application and characteristics of the pulse. The timing is adjusted    automatically so that the pulses are delivered in the first half of    the inflation phase based on the prevailing duration of the    inflation phase obtaining in previous breaths. The shape of the    applied pulse is also adjusted automatically to result in a    reasonably flat segment in Paw and flow during the pulse near T₁    (see FIG. 1). The information produced by the event processor is    conveyed to the pulse generator (DAC-08, FIG. 3) which generates a    pulse of about 0.2 second duration or any other convenient duration.    A pulse invertor and gating circuit (LM660 OP-MPS and CD4052    analogue multiplexor, FIG. 3) is used to produce either a positive    pulse or negative pulse.-   (4) A subprogram that causes the values of flow, volume, and airway    pressure, sampled at about 6 msec or other convenient time interval,    to be stored in data memory over the entire period of inspiratory    flow in breaths receiving pulses.-   (5) A subprogram that scans the above data to determine the time at    which peak Paw occurred prior to the negative deflection (T_(O)), a    time about 100 msec or other convenient time interval prior to T_(O)    (T⁻¹), the time of occurrence of minimum Paw during the pulse (T₁)    and the time of highest Paw in the post-pulse phase (T₂).-   (6) A subprogram to tabulate values of Paw, {dot over (V)}, and    volume at these four time points for each pulsed breath.-   (7) A subprogram that deletes data points that fall outside the    normal variability of the data. This subprogram also identifies    breaths subjected to a pulse perturbation where certain criteria are    not met. Data related to these observations are deleted from the    tables.-   (8) A program that determines the amplitude of pulses to be    delivered. This is an iterative program. The pulse generator is    initially instructed to deliver negative pulses of small amplitude.    The decrease in flow during these pulses is noted. If the trough of    the flow (i.e. {dot over (V)} @T_(I)) is above about 0.2 l/sec or    other convenient threshold value, the amplitude of the next negative    pulse is increased and the trough in flow is again noted.    Progressive increase in the amplitude of consecutive negative pulses    continues until the trough falls at approximately 0.2 l/sec or other    selected threshold value. The amplitude of the pulses is then kept    constant but the trough flow is monitored each time. Should the    trough rise above 0.2 l/sec or other selected threshold value and    remain elevated for a number of pulses, the amplitude of the pulse    is increased again. Conversely if the trough results in zero flow    with resetting of respiratory cycle, the amplitude of the pulse is    decreased. The intent of this subprogram is to maintain the    amplitude of the negative pulses such that the trough in flow during    the negative pulses is close to, but not zero.-   (9) A subprogram that causes early data to be deleted as new data    are acquired, leaving only the results of a specified number of    pulses (e.g. last 20 pulses) in the tables.-   (10) A statistical subprogram to calculate the values of respiratory    system resistance (R) from equations 8, 8a, 8 inter, 8a inter, 8    (bextra) and 8a (bextra) described above. These derivations may be    obtained from the average values of flow, volume and airway pressure    tabulated for negative or positive pulses, as described in detail    above.-   (11) A function which results in the display of the results of    determined resistance (R) on an LCD or other display, as required.

Whereas in the embodiment described above, a free-standing microcontroller is illustrated, the same functions performed by this microcontroller can be incorporated into a resident computer within theventilator by suitable programme. It is also recognized that theapplication of this technology is not limited to use with the specificpiston-based PAV delivery system used in the above preferred embodiment.All commercial ventilators suitable for use in the Intensive Care Unitare capable of providing outputs related to flow and airway pressure andthose commercial products which have PAV delivery capabilitiesnecessarily have circuitry or micro controllers that execute the PAValgorithm and which can be interfaced with the automated mechanicsmeasurement system provided herein. Understandably, the system describedabove may have to be changed appropriately to adapt to differentfeatures in various PAV delivery system, but any such modificationsrequired would be well within the skill of anyone experienced in theart. It is also evident that microprocessors and electronic accessoriesother than those described in the preferred embodiment can be utilizedto accomplish the same objectives.

It is also recognized that modifications to the algorithms describedabove with respect to the preferred embodiment are possible. Theseinclude, but are not limited to, the following:

-   -   1) Using pulse durations that are smaller or longer than 200        milliseconds.    -   2) Using positive pressure pulses instead of negative pressure        pulses or use of both positive and negative pulses.    -   3) The use of complex pulse forms, for example but not limited        to, bi-phasic pulses instead of unimodel pulses.    -   4) More than one pulse is applied during a given breath.    -   5) Where transient increases or decreases in applied pressure        for the sake of determining resistance are produced by        transiently changing the gain of the PAV assist.    -   6) Where transient perturbations in pressure and flow are        produced by a mechanical system independent of the ventilator        itself and incorporated in the external tubing.    -   7) Where transient perturbations in pressure and flow, for the        sake of determining resistance, are applied during modes other        than PAV, including volume cycled assist, CPAP mode, pressure        support ventilation or airway pressure release ventilation,        whereby perturbations are produced by superimposing positive        and/or negative transients to the usual control signal of the        relevant mode.    -   8) Provision to store the resistance results over extended        periods of time to be made available for later display to        provide time-related trends in such relationships.    -   9) Provision of appropriate circuitry or digital means to effect        automatic changes in the magnitude of flow assist in the PAV        mode or assist level in other modes, as the resistance values        change (i.e. closed loop control of assist level).    -   10) Where resistance is computed from values obtained from        single pulses as opposed to averages of values obtained from        several pulses.    -   11) Where the behavior of Pmus during the pulse is calculated by        interpolation between values at T_(O) and values beyond T₁ (as        per equation 8 inter and 8a inter) as opposed to the preferred        method of extrapolation of data between T⁻¹ and T_(O) (as per        equation 8 and 8a).    -   12) Where the behavior of Pmus during the pulse is calculated by        backextrapolation of values occurring between T_(I) and a point        beyond T_(I), as per equations 8 (bextra) and 8a (bextra).    -   13) Where resistance is calculated both by the extrapolation        technique (equation 8 or 8a) and interpolation technique        (equation 8 inter and 8a inter) and the result is given as an        average, or derivative, of the results of the two methods of        calculation.    -   14) Where flow is maintained nearly constant for a period of        time beyond T₁ instead of allowing it to rise again.    -   15) When the assist is terminated immediately after T₁.

SUMMARY OF DISCLOSURE

In summary of this disclosure, the present invention provides method andapparatus to determine respiratory resistance (R) during assistedventilation of a patent in a unique and simplified manner. Modificationsare possible within the scope of the invention.

1. A method of determining respiratory system resistance (R) in apatient receiving gas from a ventilatory assist device (ventilator),comprising: estimating the flow rate ({dot over (V)}) and volume (V) ofgas received by the patient from the ventilator; estimating pressurenear the airway of the patient (Paw); generating a signal that resultsin a step decrease (negative pulse) in the pressure and/or flow outputof the ventilator during selected inflation cycles; measuring Paw, {dotover (V)} and V at a point (T_(O)) near the beginning of the pulse(Paw_(O), {dot over (V)}, V_(O)), at a point (T_(I)) near the trough ofthe negative pulse (Paw₁, {dot over (V)}₁, V₁), and at a point (T⁻¹)preceding T_(O) but after the onset of inspiratory effort (Paw⁻¹, {dotover (V)}⁻¹, V⁻¹); and calculating the value of resistance (R) from thedifferences between Paw, {dot over (V)} and V at T_(O) and at T_(I) andwhere the change in patient generated pressure (Pmus) in the intervalT_(O)→T_(I) (ΔPmus(T_(O)→T_(I))) is estimated, by extrapolation, fromthe differences between Paw, {dot over (V)} and V at T_(O) and at T⁻¹,in accordance with equation (8), as follows:R=[(Paw _(O) −Paw ₁+(Δt/Δt _(−I))(Paw _(O) −Paw ⁻¹))−E(V _(O) −V₁+(Δt/Δt _(−I))(V _(O) −V ⁻¹))]/(V _(O) −V ₁+(Δt/Δt _(−I))({dot over(V)} _(O) −{dot over (V)} ⁻¹))  (8).
 2. The method of claim 1 whereinthe estimating by extrapolation step is modified by estimatingΔPmus(T_(O)→T_(I)) from the differences between Paw, {dot over (V)} andV values obtained at two time points preceding T_(O), as opposed toT_(O) and a single preceding time point (T⁻¹).
 3. The method of claim 1wherein the estimating by extrapolation step is modified by estimatingΔPmus(T_(O)→T_(I)) using regression coefficients obtained fromregression analysis of Paw, {dot over (V)} and V values measured atmultiple (>2) points preceding the pulse.
 4. The method of claim 1wherein the estimating step to estimate ΔPmus(T_(O)→T_(I)) is modifiedby estimating by interpolation, from the differences between Paw, {dotover (V)} and V values obtained at T_(O) and at a second point (T₂)beyond T_(I), in addition to, or instead of, extrapolation ofdifferences between values at T_(O) and T⁻¹.
 5. The method of claim 1 or3 wherein ΔPmus(T_(O)→T_(I)) is estimated by back extrapolation ofvalues obtained beyond T_(I).
 6. The method of claim 1 wherein thesingle R value in equation (8) is replaced by mathematical functionsthat allow for non-linear pressure-flow relations.
 7. The method ofclaim 6, wherein said mathematical function is given by R=K_(I)+K₂{dotover (V)}, wherein K1 and K2 are the coefficients defining thenon-linear pressure-flow relation.
 8. The method of claim 7, wherein K₂is replaced by a known or assumed K₂ value of the endotracheal tube ofthe patient.
 9. The method of claim 6 or 7 wherein the coefficientsdefining the non-linear pressure-flow relation (K₁, K₂) are obtained byregression analysis performed on the results of two or more pulsesapplied in separate breaths.
 10. The method of claim 1, wherein adefault elastance value (E) is used in lieu of an actually measuredelastance value for the sake of computing differences in elastic recoilpressure between T_(O), T_(I) and T_(−I) in equation (8).
 11. The methodof claim 1, wherein positive pulses are delivered instead of, or inaddition to, negative pulses and the T_(I) values of Paw, {dot over (V)}and V are measured at or near peak Paw or flow of the positive pulse.12. The method of claim 1, including automatically adjusting theamplitude of the pulse depending on the response to previous pulses. 13.The method of claim 1, including automatically adjusting the timing ofpulse application during the inflation phase.
 14. The method of claim 1,including automatically adjusting the shape of the pulse to insure thepresence of a flat segment in the Paw/flow signal during the pulse foruse in measuring T_(I) values.
 15. The method of claim 1, wherein thepulses are delivered at random intervals.
 16. The method of claim 1including user selecting one or more pulse characteristics.
 17. Themethod of claim 1, wherein the resistance results (R) are reported asaverages of the results of several pulses.
 18. The method of claim 1,wherein resistance results (R) are stored over time to permit thedisplay of time dependent trends.
 19. The method of claim 1 includingidentifying pulses with results that fall outside the normal variabilityof the data as determined from several data sampling and excluding theresults of these pulses from analysis and determination of R.
 20. Themethod of claim 1 including deleting early data as new data are acquiredand reporting the results of the determination of R for a specifiednumber of pulses.
 21. The method of claim 20 wherein the specifiednumber of pulses is selected, either by a user or, in the absence ofuser input, as a default value (e.g. 20).
 22. The method as claimed inclaim 1, wherein the step decrease or increase in Paw or {dot over (V)}is produced by an electromechanical system attached to the externaltubing of the ventilator as opposed to directly interfacing with theventilator control system.
 23. The method of claim 1, wherein theresults of the resistance values are used in closed loop control of anassist level provided by the ventilator.
 24. An apparatus whichinterfaces with ventilatory assist devices (ventilators) and whichdetermines respiratory system resistance (R), comprising: a flowmeter,with associated electronic circuitry, that estimates the flow rate ({dotover (V)}) and volume (V) of gas received by a patient; a pressuresensor that estimates pressure near the airway of the patient (Paw); andelectronic circuitry which receives the Paw, {dot over (V)} and Vsignals from above mentioned circuitry and which is also connected to acontrol system of the ventilator, comprising: circuitry that generatesan output that results in a step decrease (negative pulse) in thepressure and/or flow output of the ventilator during selected inflationcycles; circuitry that measures Paw, {dot over (V)} and V at a point(T_(O)) near the beginning of the pulse (Paw_(O), {dot over (V)}_(O),V_(O)), at a point (T_(I)) near the trough of the negative pulse (Paw₁,{dot over (V)}_(I), V₁), and at a point (T⁻¹) preceding T_(O) but afterthe onset of inspiratory effort (Paw⁻¹, {dot over (V)}_(−I), V⁻¹);circuitry to calculate the value of resistance (R) from the differencesbetween Paw, {dot over (V)} and V at T_(O) and at T_(I) and where thechange in patient generated pressure (Pmus) in the intervalT_(O)→T_(I)(ΔPmus(T_(O)→T_(I))) is estimated, by extrapolation, from thedifferences between Paw, {dot over (V)} and V at T_(O) and at T⁻¹, inaccordance with equation (8), as follows:R=[(Paw _(O) −Paw ₁+(Δt/Δt ⁻¹)(Paw _(O) −Paw ⁻¹))−E(V _(O) −V ₁+(Δt/Δt⁻¹)(V _(O) −V ⁻¹))]/({dot over (V)} _(O) −{dot over (V)} ₁+(Δt/Δt⁻¹)({dot over (V)} _(O) −{dot over (V)} ⁻¹))  (8).
 25. The apparatus ofclaim 24 wherein ΔPmus(T_(O)→T_(I)) is estimated, by extrapolation, fromthe differences between Paw, {dot over (V)} and V values obtained at twotime points preceding T_(O), as opposed to T_(O) and a single precedingtime point.
 26. The apparatus of claim 24 wherein ΔPmus(T_(O)→T_(I)) isestimated, by extrapolation, using regression coefficients obtained fromregression analysis of Paw, {dot over (V)} and V values measured atmultiple (>2) points preceding the pulse.
 27. The apparatus of claim 24wherein ΔPmus(T_(O)→T_(I)) is estimated, by interpolation, from thedifferences between Paw, {dot over (V)} and V values obtained at T_(O)and at a second point (T₂) beyond T_(I) in addition to, or instead of,as opposed to extrapolation of differences between values at T_(O) andT⁻¹.
 28. The apparatus of claim 24 or 26 wherein ΔPmus(T_(O)→T_(I)) isestimated by back entrapolation of values obtained beyond T_(I).
 29. Theapparatus of claim 24 wherein the single R value in the equations isreplaced by mathematical functions that allow for non-linearpressure-flow relations.
 30. The apparatus of claim 29 wherein saidmathematical function is given by R=K_(I)+, wherein K1 and K2 arecoefficients defining the non-linear pressure-flow relation.
 31. Theapparatus of claim 30 wherein K₂ is replaced by a known or assumed K₂value of the endotracheal tube of the patient.
 32. The apparatus ofclaim 29 or 30 wherein the coefficients defining the non-linearpressure-flow relation (K₁, K₂) are obtained by regression analysisperformed on the results of two or more pulses applied in separatebreaths.
 33. The apparatus of claim 24 wherein a default elastance value(E) is used in lieu of an actually measured elastance value for the sakeof computing differences in elastic recoil pressure between T_(O), T_(I)and T_(−I).
 34. The apparatus of claim 24 wherein positive pulses aredelivered instead of, or in addition to, negative pulses and the T_(I)values of Paw, {dot over (V)} and V are measured at or near peak Paw orflow of the positive pulse.
 35. The apparatus of claim 24 includingalgorithms to automatically adjust the amplitude of the pulse dependingon response to previous pulses.
 36. The apparatus of claim 24 includingalgorithms which automatically adjust the timing of pulse applicationduring the inflation phase.
 37. The apparatus of claim 24 includingalgorithms that automatically adjust the shape of the pulse to insurethe presence of a flat segment in the Paw/flow signal during the pulsefor use in measuring T_(I) values.
 38. The apparatus of claim 24 whereinthe pulses are delivered at random intervals.
 39. The apparatus of claim24 including a user interface that permits the user to select one ormore pulse characteristics.
 40. The apparatus of claim 24 wherein theresistance results are reported as averages of the results of severalpulses.
 41. The apparatus of claim 24 wherein resistance results arestored over minutes, hours or days to permit the display of timedependent trends.
 42. The apparatus of claim 24 including algorithmswhich identify pulses with results that fall outside the normalvariability of the data and which exclude the results of these pulsesfrom analysis.
 43. The apparatus of claim 24 including algorithms whichdelete early data as new data are acquired and which report the resultsof a specified number of pulses.
 44. The apparatus of claim 43 whereinthe specified number of pulses is selected by user or, in absence ofuser input, as a default value (e.g. 20).
 45. The apparatus of claim 24wherein some or all necessary components are incorporated within themain body of the ventilator.
 46. The apparatus of claim 24 wherein thestep decrease or increase in Paw or {dot over (V)} is produced by anelectromechanical system attached to the external tubing of theventilator as opposed to directly interfacing with the ventilatorcontrol system.
 47. The apparatus of claim 24 wherein the results of theresistance values are used in closed loop control of the assist levelprovided by the ventilator.